This invention relates generally to systems and methods for imaging tissue using magnetic resonance imaging (MRI) techniques, and more particularly to obtaining high accuracy temperature measurements of a tissue mass by compensating for errors during the acquisition of MR images used for temperature measurement by applying a T1-based correction to the MR images.
In the past, ultrasound imaging has been used for both generating and positioning focused ultrasound waves on tissue to be destroyed. It was soon found that the MR images were much better suited for such guidance, since MR provides excellent images of tissue, and is not limited to “windows” that excludes bones, for example. Further, MRI system images are useful not only for guiding the actual surgical procedures, but also for planning the surgical procedures. In fact, a tumor is often much more visible in an MR image than as seen in actual surgery. In actual surgery, the tumor and normal tissues often look similar. In addition, during the surgery, the tumor can be obscured by blood. Presently, patients are first scanned in an MRI system to locate the tumor and plan a safe trajectory between the entry of the ultrasound radiation and the target points.
Certain types of tissues, such as cancerous tumors, can be destroyed by heat. It has further been discovered that the control of heat was improved by using short pulses such that the effect of blood profusion is made negligible. The state of the art has advanced to the point where the MRI system is used not only for planning the surgery, but also during the actual destruction of the cancerous tissue. The MRI system using separate scanning sequences provides temperature information and, in addition, also provides tissue information. Thus, the actual temperature of the tissue can be ascertained using magnetic resonance imaging methods, and in addition the ablation of the tissue can be observed using the MRI system.
One conventional way to heat these tissues is by directing laser energy into the tissue using, for example, a laser source carried by a catheter. Another conventional way is to focus high intensity, ultrasonic acoustic waves into the tissue using, for example, a phased-array of piezoelectric transducers. Both of these approaches can reduce or even eliminate the need for invasive surgery to remove the tissue.
Thus, the prior art performs thermal surgery guided by MRI systems and procedures to selectively destroy tumorous tissue in the patient with localized heating, without adversely affecting healthy tissue. The heat is applied to the tumor tissue in a pulsed or oscillating fashion. The pulsed energy creates a heat focus that heats either at the tip of the optical fiber, or at the focal point of the transducer, depending on the heat source. The heated tumorous region may be imaged with the use of the MRI systems, employing a temperature sensitive MR pulse sequence to acquire a temperature “map” that is used basically to assure that the heat is being applied to the tumorous tissue and not to the surrounding healthy tissue. This is done by applying a quantity of heat that is insufficient to cause necrosis but is sufficient to raise the temperature of the heated tissue. The MRI system temperature map shows whether or not the heat is applied to the previously located tumorous tissue. The imaging system is also used in a separate scan sequence to create an image of the tissue intended to be destroyed. Using the imaging system in the prior art, the operator of the apparatus adjusts the placement of the radiation on the site of the tissue to be destroyed. The MR image of the tumor acquired in the separate scan determines in real time if necrosis is occurring and effectively ablating the tumorous tissue. However, the monitoring and guiding are provided using separate two-dimensional scan sequences.
When laser beams are used as the heat source, then mechanical means are used to position the optical fiber carrying the laser beam to the diseased tissue to accurately position it within the diseased tissue. When ultrasound beams are used as the heat source, then the focus of the ultrasound beams can be positioned either by mechanically moving the ultrasound generator to move its focal point, or a phased array ultrasound system can be used to manipulate the position of the focal point so that it is within the diseased tissue that is to be destroyed by heating.
Of critical importance to the process is verifying that a sufficient temperature was reached during each application of ultrasonic energy to kill the target tissue structure, or portion thereof being heated. This can be done by measuring the temperature change (rise) of the portion of the tissue structure being heated during the heating process using MR imaging techniques.
There are several MRI methods that have been used in the past for measuring temperatures using well-known MRI parameters, such as the spin-lattice relaxation time T1, the time to repeat (TR), the time to echo TE, and the flip angle. For example, temperature maps can be generated based on such procedures that provide T1 derived images evaluated with fast spoiled gradient echo sequences applied during the actual thermal therapy exposure. The parameters used are to some degree based on the tissue type and the precise evaluation of the behavior due to physiological or metabolic changes in the tissue during thermal therapy exposure. For example, TE, TR and the flip angle of the spoiled gradient echo have been used for localizing the low-temperature elevation induced by a focused ultrasound beam during both the planning and treatment.
Another known method of measuring temperature change using MR techniques exploits the temperature dependence of the proton resonant frequency (PRF) in water. The temperature dependence of the PRF is primarily due to temperature-induced induced rupture, stretching, or bending of the hydrogen bonds in water. The temperature dependence of pure water is 0.0107 ppm/° C., and the temperature dependence of other water-based tissues is close to this value.
Because of a non-homogenous magnetic field within the MR image, absolute PRF measurements are not possible. Instead, changes in PRF are measured by first taking a MR image before the delivery of heat, and subtracting this baseline value from subsequent measurements. Notably, the total imaging time must be kept relatively short for the baseline value to remain relatively stable, since drifts in the magnetic field can occur over time.
The temperature-induced changes in PRF are then estimated by measuring changes in phase of the MR signal, or proton resonance frequency shift (PRFS) in certain MR imaging sequences. Unfortunately, movements of the object being imaged or dynamic changes in magnetic field intensity and/or uniformity during the MR imaging process (i.e., between successively acquired images) can also induce phase shifts, which can be misinterpreted as a temperature-induced phase shift. As a result, a given phase shift might be attributable to any one of temperature change, motion, changes in magnetic field, or some combination thereof, between acquired images. This ambiguity makes it difficult to determine the tissue mass temperature change relying only on the MR signals from the tissue mass being heated. Such motion and non-uniform field “distortion” may have differing sources acting alone or in combination, such as patient motion, breathing or other dynamic organ movement (e.g., the heart), or blood moving through a blood vessel located in or adjacent to the target tissue region.
Another method is known as self-referencing or reference-less thermometry. In this method, tissue outside the therapeutic zone (tissue being heated) is used to measure the PRFS. Assuming this tissue does not heat, then the PRFS in these regions is due only to changes in magnetic field. Unfortunately, the temperature is these regions often do change, rendering the method ineffective.
Another method being developed is to use nearby fat tissue as a reference. Fat tissue does not experience a PRFS with temperature change, so that any temperature-induced changes in PRF are due to background field changes. The limitation here is that one may not have enough fatty tissues around to perform this correction.
In view of the foregoing, the problem of performing accurate measurement of temperature changes in tissue in an MRI system remains one of the biggest barriers to widespread use of the PRF method in temperature imaging. Therefore, it would be desirable to provide a system and method that accurately measures the temperature changes in tissue in an MRI system.